PDGFB-targeted functional MRI nanoswitch for activatable T1–T2 dual-modal ultra-sensitive diagnosis of cancer

As one of the most significant imaging modalities currently available, magnetic resonance imaging (MRI) has been extensively utilized for clinically accurate cancer diagnosis. However, low signal-to-noise ratio (SNR) and low specificity for tumors continue to pose significant challenges. Inspired by the distance-dependent magnetic resonance tuning (MRET) phenomenon, the tumor microenvironment (TME)-activated off–on T1–T2 dual-mode MRI nanoswitch is presented in the current study to realize the sensitive early diagnosis of tumors. The tumor-specific nanoswitch is designed and manufactured on the basis of PDGFB-conjugating ferroferric oxide coated by Mn-doped silica (PDGFB-FMS), which can be degraded under the high-concentration GSH and low pH in TME to activate the T1–T2 dual-mode MRI signals. The tumor-specific off–on dual-mode MRI nanoswitch can significantly improve the SNR and is used successfully for the accurate diagnosis of early-stage tumors, particularly for orthotopic prostate cancer. In addition, the systemic delivery of the nanoswitch did not cause blood or tissue damage, and it can be excreted out of the body in a timely manner, demonstrating excellent biosafety. Overall, the strategy is a significant step in the direction of designing off–on dual-mode MRI nanoprobes to improve imaging accuracy, which opens up new avenues for the development of new MRI probes. Supplementary Information The online version contains supplementary material available at 10.1186/s12951-023-01769-7.


Introduction
In situ accurate and early diagnosis of cancer is essential for selecting appropriate treatment options and determining prognosis without invasive procedures such as tissue biopsies or surgical excision [1][2][3]. In recent years, clinical molecular imaging devices, such as X-ray computed tomography (CT), optical imaging, positron emission tomography (PET), and single-photon emission computed tomography (SPECT) techniques, have been utilized extensively to analyze and diagnose tumors in vivo [4,5]. However, due to the superior soft-tissue contrast, deep tissue penetration, high spatial resolution, and lack of ionizing radiation exposure, etc., magnetic resonance imaging (MRI) can provide anatomical and functional information in diagnosing clinical diseases; consequently, it has demonstrated broader application prospects in the diagnosis of tumors [6][7][8]. Accordingly, contrast agents, such as T 1 or T 2 contrast nanomaterials, have been developed for MRI over the past few decades [9][10][11]. However, single-mode MR contrast agents (CAs) have inherent drawbacks, such as a short blood circulation time, magnetic susceptibility artifacts, and a clinical accuracy of tumor diagnosis that is not always met [12][13][14]. Therefore, it is of the utmost importance to design versatile contrast nanoprobes capable of overcoming limitations to achieve accurate diagnostic results.
The concurrent use of T 1 -T 2 dual-mode imaging modalities has been reported for the cross-validation of acquired image data [15,16]. Accordingly, this strategy combines the advantages of two single-mode imaging techniques to provide extremely precise image information. Based on the distance-dependent magnetic resonance tuning (MRET) phenomenon, however, when the enhancer is located close to the quencher, the electron spin fluctuation rate of the enhancer decreases, preventing it from accelerating and relaxing water protons, thereby resulting in a quenched longitudinal relaxivity (r 1 ) [17][18][19]. Additionally, Due to direct contact between the two agents, the magnetic field derived from the superparamagnetic T 2 CA disrupts the relaxation process of the paramagnetic T 1 CA, resulting in the attenuation of the T 1 signal and a significant decline in the T 1 -T 2 contrast enhancement [20]. Conversely, "always on" MRI contrast agents frequently disregard the specific interaction with tumor tissues exhibiting weak acidity and high GSH concentration, resulting in a low signal-to-noise ratio (SNR) [21][22][23]. Thus, the development of a smart stimuli-activated MRI off-on nanoswitch [24,25] to meet the practical requirements for early, tumor-specific diagnosis, with high SNR and off-on properties, presents a significant challenge.
In this study, inspired by the MRET principle, a novel tumor microenvironment (TME)-activated T 1 -and T 2 -dual-mode MRI nanoswitch was designed using platelet-derived growth factor (PDGFB)-conjugated ferroferric oxide coated by Mn-doped silica (FMS) yolk-shell nanostructures (PDGFB-FMS) (Scheme 1), which display a distinct "off-on" T 1 -T 2 dual-modal synergistic imaging with an ultrafast response, high sensitivity and specificity toward TME. Under normal tissue conditions, this nanostructure for dual-mode imaging is so stable that the T 1 and T 2 MR signals initially display an "off" state. However, when the tumor is targeted, the weakly acidic and high GSH of TME disrupts the PDGFB-FMS, causing its nanostructure to collapse, resulting in the rapid release of Mn 2+ ions, which separates from the Fe 3 O 4 magnetic core and activates the dual-mode of the T 1 and T 2 MRI signals. In addition, the conjugation of PDGFB cycle peptides makes the nanoswitch recognize the tumor site specifically [26][27][28][29][30][31][32], allowing the nanoswitch to be effectively internalized by tumor cells and accumulate preferentially at tumor sites. This smart activatable dualmode MRI nanoswitch not only responds specifically to TME but also exhibits no signal (or a very weak signal) in normal tissues, which can reduce background noise and improve SNR. We, therefore, anticipate that the dualmode nanoprobes will play a major role in the development of a variety of diagnostic nanoplatforms.

Synthesis of FMS yolk-shell nanoswitch
The preparation of FMS consisted of three steps, as illustrated in Scheme 1a. We first synthesized Fe 3 O 4 nanoclusters via the solvent thermal decomposition method using ferric acetylacetonate as the iron source and sodium acrylate as the template. As depicted in Fig. 1a, the Fe 3 O 4 nanoclusters were monodispersed and uniformly spherical with a size of approximately 60 nm. Moreover, the hydrodynamic diameter of the Fe 3 O 4 nanocluster was approximately 60 ± 19.7 nm, which is comparable to the size obtained from transmission electron microscopy (TEM) studies, indicating that the Fe 3 O 4 nanocluster had an excellent colloidal performance. In addition, the high-resolution TEM (HRTEM) image reveals that the Fe 3 O 4 nanoclusters were highly crystalline, with a lattice fringe distance of ~ 0.253 nm, which corresponds to the (311) atomic plane of magnetite (Fig. 1b). Subsequently, the well-dispersed spherical Fe 3 O 4 @SiO 2 (FS) core-shell nanoparticles with a shell thickness of ~ 35 ± 16.0 nm were formed by using TEOS as the silica source to coat the Fe 3 O 4 under heating and alkaline conditions (Fig. 1c). Finally, a specific "ammonia-assisted Mn 2+ etching" strategy was used to etch the outer silica shell to form Mndoped silica-coated ferroferric oxide nanoparticles (FMS) via the Ostwald ripening mechanism [33]. Compared to untreated FS nanoparticles, the representative TEM image revealed well-defined yolk-shell nanostructures with a larger size (~ 220 ± 100.8 nm) (Fig. 1d). In addition, the FMS exhibited a porous structure that was loose and rough, providing an abundance of active sites for proton exchange. Furthermore, the increasing hydrodynamic size of nanoparticles from FS to FMS also confirmed the TEM observations (Additional file 1: Fig. S1). The highangle annular dark field scanning TEM (HAADF-STEM) image of FS revealed the composition and nanostructure of the core-shell, but FMS further illustated the successful doping of Mn ions and nanostructure changes from core-shell to yolk-shell (Fig. 1e, f ). Moreover, the energydispersive X-ray (EDX) spectra showed that all expected elements (Fe, Si, Mn, and O) were detected and their relative positions in yolk-shell nanostructures are well matched (Additional file 1: Fig. S2g-i).
In addition, X-ray diffraction (XRD) spectra of Fe 3 O 4 revealed the characteristic spinel structure of magnetite ( Fig. 1g) [34]. Upon growth of the SiO 2 shell layer, a characteristic broad peak of amorphous silica is produced, whereas it nearly disappears after Mn 2+ ions etch the FMS. This result indicates that the silica shell is successfully coated on the surface of Fe 3 O 4 nanoparticles, and it is effectively etched by Mn 2+ in the subsequent process. After Mn 2+ etching, significant increases in specific surface area (342.92 m 2 /g) and porosity are observed for FMS nanostructures with a pore size distribution dominated by 2.0 nm (Fig. 1h). As is well known, the weight loss of the thermogravimetric (TG) curves below 200 °C is attributable to the loss of water molecules, including physically adsorbed and bound water. Accordingly, the TG analysis suggests that FMS adsorbs significantly more water (7.98%) than FS (6.58%) and Fe 3 O 4 (2.35%) (Additional file 1: Fig. S2a). These findings indicate that FMS has a strong affinity for water molecules and facilitates the interaction with the proton, which can improve the relaxation properties. In the Fourier transform infrared (FT-IR) spectrum of Fe 3 O 4 , the peak of the Fe-O bond appears at 570 cm −1 [34], while the typical peaks of the -C=O and C-H bonds appear at 1635 cm −1 and 2920 cm -1 , respectively (Additional file 1: Fig. S2e). In addition, the FT-IR spectrum of FS exhibits a distinct Si-O peak at 1100 cm −1 [35], indicating that silica has been successfully coated on the surface of  Fig. S2c), further confirming the transformation of Si from a dense silica layer to a loose structure. The hydrodynamic size of PDGFB-FMS had no significant variation in blood, phosphate-buffered saline (PBS), and 10% fetal bovine serum (FBS) within 3 days, demonstating excellent stability (Additional file 1: Fig. S3).

TME-responsive "off-on" T 1 -T 2 imaging of FMS nanoswitch
We also analyzed the degradation of FMS under GSH and weak acid conditions. In the presence of weak acidity and GSH, Fig. 2a, b illustrates the significant effects of FMS on the hydrodynamic size distribution and morphology of corresponding particles. Most FMS particles disassemble from regular spherical structures into scattered dots over time. In contrast, no discernible change is observed for FMS in the absence of GSH and under neutral conditions. In addition, the degree of dissociation of FMS increased significantly with increasing GSH concentration and decreasing pH, demonstrating excellent pH-and GSH-responsive degradability (Additional file 1: Fig. S4). To gain more insight into the evolution process, the release of Mn 2+ ions from FMS with the change of GSH concentration and pH value was investigated.
Accordingly, the results confirm that the behavior of Mn 2+ ion release from FMS is pH-and GSH-dependent (Fig. 2c, d). These results indicate that FMS can be responsively dissociated in TME with an abundance of GSH and an acidic environment by breaking the Mn-O bond and the subsequent collapse of the FMS backbone within one hour (Fig. 2c, d). Based on the excellent degradation of FMS triggered by TME, we further investigated the off-on switchable properties of the dual-mode T 1 -T 2 signal in FMS. Under physiological conditions (in pH 7.4 PBS), both T 1 -and T 2 -weighted imaging (T 1 WI and T 2 WI, respectively) exhibited an apparent quenching phenomenon (Fig. 2e, g). In addition, the corresponding T 1 relaxivity (r 1 ) and T 2 relaxivity (r 2 ) were measured to be 5.1 and 126.46 mM −1 s −1 , respectively (Fig. 2f, h). When exposed to GSH (20 mM) and pH 4.5, the FMS nanostructure decomposed rapidly into Fe 3 O 4 and Mn 2+ ions, resulting in a rapid recovery of T 1 and T 2 signals as the distance between the Mn 2+ ions and Fe 3 O 4 magnetic core increased (Fig. 2e, g). After treatment with PBS containing 20 mM GSH and pH 4.5, the calculated r 1 and r 2 values were found to be 11.06 and 304.22 mM −1 s −1 , respectively (Fig. 2f, h). In addition, as depicted in Additional file 1: Figs. S5, S6, the "off-on" characteristics of dual-mode T 1 -T 2 nanoprobes were highly dependent on GSH concentrations and pH values. These results, therefore, support our hypothesis that low pH and high GSH concentrations triggered the release of Mn 2+ , which reduces the mutual magnetic resonance interference between Mn 2+ and Fe 3 O 4 and activates the T 1 -T 2 MRI signals (Additional file 1: Tables S1 and S2).

Quenching mechanism of T 1 and T 2 imaging
The T 1 -T 2 dual-quenching mechanism of FMS was also investigated methodically. Accordingly, Mn 2+ ions with a single T 1 contrast and Fe 3 O 4 with a single T 2 contrast were used as controls (Additional file 1: Fig.  S7a and Fig. 3). For the T 1 quenching effect, we proposed that the strong local magnetic field originating from the large magnetic moment of the Fe 3 O 4 magnetic core would interfere with the spin-lattice relaxation process of Mn-doped silica shell with much weaker paramagnetism surrounding it, thereby quenching the T 1 signal of paramagnetic Mn 2+ . Accordingly, the M-H curves of FMS exhibited a lower saturation magnetization (25 emu/g) at 300 K in comparison to Fe 3 O 4 (112.4 emu/g). In addition, the magnetic field generated by the superparamagnetic Fe 3 O 4 core significantly disrupts the relaxation process of the paramagnetic Mn-doped silica shell, resulting in the quenching of the T 1 signal during the etching procedure (Fig. 3a). Meanwhile, the grayscale and corresponding pseudo-color T 2 WI of Fe 3 O 4 , FS, and FMS exhibited significant T 2 quenching effects (Fig. 3b, c), and their T 2 relaxivities (r 2 ) were 317.4 mM −1 s −1 , 203.2 mM −1 s −1 , and 126.5 mM −1 s −1 , respectively (Fig. 3d). Therefore, T 1 and T 2 MRI signals can be interfered with when Fe 3 O 4 and Mn-doped silica are in direct contact. To further validate the MRET hypothesis, we synthesized a series of FMS (FMS-1, FMS-2, FMS-3, FMS-4, FMS-5) with different separation distances between the paramagnetic Mn 2+ and the Fe 3 O 4 magnetic core by varying the thickness of the silica separation layer from 40 ± 2.6 to 0 ± 1.2 nm (Fig. 3e). Notably, the hydrodynamic size of FMS increased gradually as the etching process progressed (Additional file 1: Fig. S7b), while simultaneously increasing the specific surface area and porosity (Additional file 1: Fig. S8), facilitating the energy exchange with water molecules. The nanoswitch showed obvious quenching in T 1 and T 2 signals with the Mn 2+ doping ratio from 33 to 88% (Additional file 1: Fig. S9). This was because that high ratio of Mn-doped silica significantly reduced the distance between Fe 3 O 4 core and Mn 2+ , resulting in the increase of quenching effect in T 1 and T 2 imaging. As shown in Fig. 3f, the T 1 MRI signal of the Mn-doped silica shell in FMS decreases as the Mn approaches the Fe 3 O 4 core. Furthermore, as the separation distance decreases from 40 to 30, 20, 10, and 0 nm, the r 1 values of the Mn-doped silica shell decrease from 9.7  (Fig. 3g and Additional file 1: Fig. S7c), which is lower than that of free Mn 2+ (r 1 = 15.9 mM −1 s −1 ) (Additional file 1: Fig.  S7a). Concurrently, a similar T 2 MRI signal quenching phenomenon was observed in T 2 WI and corresponding pseudo-color MR images (Fig. 3f ). As the etching depth increases, the r 2 value of Fe 3 O 4 decreases from 190.85 to 177.66, 149.15, 134.67, and 126.5 mM −1 s −1 (Fig. 3h and Additional file 1: Fig. S7d). In addition, the saturated magnetization of FMS decreased gradually with increasing Mn content, which could be attributed to the disruption of Fe 3 O 4 magnetic moments induced by paramagnetic Mn 2+ (Fig. 3i). Thus, the aforementioned results demonstrate that the T 1 and T 2 dualquenching effect of FMS is inversely proportional to the separation distance, which can be precisely controlled by the Mn etching process.

Evaluation of internalization and cell MRI of PDGFB-FMS
The hypoxic TME stimulates the overexpression of PDGF-β receptors in breast cancer cells and prostate cancer cells [36,37]. To obtain tumor-specific delivery, the PDGFB targeting ligand is, therefore, conjugated to FMS to create a PDGFB-FMS nanoswitch that targets tumors. The internalization efficacy of a nanoprobe is essential for evaluating its targeting capability and contrast performance. Using confocal laser scanning microscopy (CLSM) and inductively coupled plasma mass spectrometry (ICP-MS) analysis, the uptake of FMS and PDGFB-FMS was, therefore, investigated in 4T1 cells and PC-3 cells based on this information (Fig. 4a-d, and Additional file 1: Figs. S10, S11). The results of CLSM indicated that the internalization of FMS and PDGFB-FMS is concentration-and time-dependent. In addition, PDGFB-FMS treated cells exhibited a stronger green fluorescence than FMS-treated cells, confirming their superior targeting ability to breast cancer and prostate cancer. We also calculated the colocalization and particle uptake ratios corresponding to the CLSM images. We found that the differences in cell internalization for FMS and PDGFB-FMS treatments are statistically significant (Additional file 1: Fig. S12). In addition, Fe content was used as an indicator for ICP-MS analysis to determine the internalization efficacy of PDGFB-FMS. Results indicate that Fe content in 4T1 cells increased gradually with increasing time and concentration, and cells treated with PDGFB-FMS accumulated more Fe than those treated with FMS. These results are consistent with the CLSM observation, further validating the performance of PDGFB-FMS in targeting tumors.
The imaging performance of PDGFB-FMS was further investigated at the cellular level, in light of the aforementioned encouraging findings. The T 1 and T 2 -weighted signals of 4T1 cells treated with FMS and PDGFB-FMS were found to be significantly activated (Fig. 4e, f, h, i) and demonstrate a time-activation relationship. Notably, the T 1 and T 2 signal almost remained unchanged in THLE-3 cells incubated with FMS and PDGFB-FMS, which can be attributed to the lower cellular GSH level and increased internalization efficacy. To determine whether the activation of PDGFB-FMS correlates positively with intracellular GSH levels, 4T1 cells were pretreated with various concentrations of GSH and then incubated continuously with PDGFB-FMS. The results revealed that both the T 1 and T 2 signals of cells were significantly improved (Fig. 4g, j). Moreover, the corresponding relaxation rate change ΔR 1 (R 1 = 1/T 1 ) and ΔR 2 (R 2 = 1/T 2 ) increased gradually after GSH pretreatment (Additional file 1: Table S3), suggesting that the intracellular high level of GSH can be used as a stimulus to activate the dual-mode MRI contrast enhancement.

In vivo MRI studies of PDGFB-FMS
A series of in vivo MRI experiments were also designed and conducted using a mouse model bearing the 4T1 tumor to validate the activable MRI diagnosis for the tumor. FDA-approved "Magnevist" (Gd-DTPA) and "Feridex" (Fe 3 O 4 ) were selected as controls for T 1 WI and T 2 WI, respectively. Accordingly, the 4T1 tumorbearing mice were intravenously administrated using Gd-DTPA, Fe 3 O 4 , FMS, and PDGFB-FMS at the dosage of 5 mg/kg, respectively, following which, they were scanned using a 7.0 T MRI scanner, and the T 1 WI and T 2 WI of the tumor in the axial plane were obtained at various time points. As shown in Fig. 5a, the T 1 WI of tumors treated with PDGFB-FMS brightens significantly at 0.5 h post-injection (p.i.) and then darkens over time. In addition, the tumor site of the PDGFB-FMS group in T 1 WI was observed to be the clearest and brightest compared to those of FMS and Gd-DTPA. Similarly to T 1 WI, the T 2 WI of the tumor treated with PDGFB-FMS was observed to be darkest at p.i. 0.5 h (Fig. 5c), following the same pattern as T 1 WI. On the contrary, the contrast enhancement of FDA-approved Gd-DTPA and Fe 3 O 4 for T 1 WI and T 2 WI, respectively, was found to be weak. In addition, the quantitative analysis of tumor regions is performed by measuring the change in SNR (ΔSNR) before and after administration. As shown in Fig. 5b, the maximum ΔSNR of the tumor regions in the PDGFB-FMS group was greater than that of the FMS and Gd-DTPA groups, demonstrating the most pronounced contrast enhancement. In addition, T 2 WI of mice treated with FMS exhibited a low ΔSNR, whereas, PDGFB-FMS exhibited a comparable ΔSNR (Fig. 5d), which may be a result of efficient targeting. In addition, the respective T 1 and T 2 map images confirmed the aforementioned findings. As shown in Additional file 1: Fig. S13, there was a 30.7% decrease in the T 1 map value in the tumor region treated with PDGFB-FMS at p.i. 0.5 h, whereas the Gd-DTPA and FMS groups only led to a 16.5% and 16.6% decrease, respectively. Similarly, a significant decrease in the T 2 map value is observed at p.i. 0.5 h, and it follows the same pattern as the T 1 map value. The aforementioned results demonstrate that the PDGFB-FMS dualactivated nanoswitch has the desired dual-modal imaging capability for accurate tumor diagnosis.
To further illustrate GSH tuning nanoswitch, we examined the T 1 and T 2 tumor map images following PDGFB-FMS treatment in the absence and presence of GSH and GSH inhibitor (LBS). Briefly, mice carrying 4T1 tumors were intratumorally pretreated with 10 mM GSH, 2.5 mM LBS (LBS-1), or 5 mM LBS (LBS-2), respectively. The PDGFB-FMS was then administered intravenously to these 4T1 tumor-bearing mice. As anticipated, the GSHpretreated tumor exhibited lower T 1 and T 2 map values at the same PDGFB-FMS injection as compared to the non-GSH-pretreated tumor. In contrast, T 1 and T 2 map value attenuation of the tumor after LBS pretreatment was significantly inhibited, and their inhibition effect was found to be concentration-dependent (Fig. 5e-g). This phenomenon indicates that the decreased intratumoral GSH concentration caused by LBS prevented the activation of GSH-activated T 1 -T 2 signals. The findings, therefore, clearly demonstrate the clear positive correlation between PDGFB-FMS activation and intracellular GSH concentrations.
The imaging potential of the PDGFB-FMS nanoswitch for in-situ carcinoma was also investigated (Fig. 6). This study establishes the transgenic adenocarcinoma of mouse prostatic (TRAMP) model to simulate the progression of orthotopic prostate cancer [38,39]. We first assessed the cytotoxicity of PDGFB-FMS nanoswitch via cell counting kit-8 (CCK-8) assays before in vivo MRI studies. No apparent cytotoxicity against PC-3 cells was observed for PDGFB-FMS nanoswitch, indicating that PDGFB-FMS nanoswitch has good biocompatibility (Additional file 1: Fig. S14). Then, transgenic adenocarcinoma mouse prostate (TRAMP) mice were administered Gd-DTPA, Fe 3 O 4 , FMS, and PDGFB-FMS intravenously, respectively. Correspondingly, the T 1 and T 2 MR contrasts of the prostate tumor site were significantly enhanced 1 and 4 h post-injection of PDGFB-FMS, respectively. However, the modest dynamic contrast enhancement of T 1 and T 2 in the Gd-DTPA, Fe 3 O 4 , and FMS groups was visible 8 h after injection (Fig. 6a,  b, g, h). The ΔSNR was then used to quantify the MR signals of tumors at different time points (Fig. 6c, i). Accordingly, for T 1 -weighted MRI, the maximum ΔSNR of tumors treated with PDGFB-FMS reached up to 75.6 ± 8.8%, which was significantly greater than that of tumors treated with FMS (39.9 ± 6.3%) and Gd-DTPA (32.7 ± 4.5%) (Fig. 6c). Using T 2 -weighted MRI, the maximal ΔSNR of tumors treated with PDGFB-FMS, FMS, and Fe 3 O 4 is 33.3%, 25.7%, and 39.7%, respectively (Fig. 6i). These results suggest that PDGFB-FMS targets tumor sites and that T 1 and T 2 imaging are effectively illuminated by TME. Furthermore, we collected T 1 and T 2 map images to verify this conclusion (Fig. 6d, e, j, and k). At the tumor treated with PDGFB-FMS, approximately 33.3% of reduction in T 1 map value was detected, whereas FM and Gd-DTPA only led to 19.4% and 10.8% reduction in T 1 map value, respectively (Fig. 6f ). Moreover, PDGFB-FMS exhibited a significant decrease in T 2 map value (41.7% at p.i. 4 h), whereas T 2 map value changes for FMS and Fe 3 O 4 were 31.9% and 53.0%, respectively (Fig. 6l). Considering the significantly higher T 1 and T 2 ΔSNR and the sensitive "off-on" dual-mode MRI signals, the PDGFB-FMS nanoswitch demonstrates its superiority in the area of early-stage cancer detection.

In vivo biosafety assessment of PDGFB-FMS nanoswitch
As is well-known, the biosafety of a nanoagent is a crucial factor to consider when assessing its clinical application potential [40,41]. Based on this, we first evaluate the biocompatibility of PDGFB-FMS with blood using a hemolysis assay and a routine blood examination. As shown in Fig. 7a and b, the hemolysis rate of PDGFB-FMS is extremely low and negligible, indicating that PDGFB-FMS cannot damage red cells. In addition, major blood routine indicators such as WBC, RBC, PLT, HGB, and HCT (Additional file 1: Fig. S15) at p.i. PDGFB-FMS did not demonstrate any significant changes. These results demonstrate the biocompatibility of PDGFB-FMS with blood. Subsequently, we also investigated cytoxicity of PDGFB-FMS on normal cells (THLE-3, 293T, PC-12, and SH-SY5Y cells). As shown in Additional file 1: Fig.  S16a-c, the PDGFB-FMS is biocompatible with THLE-3, c ΔSNR corresponding to a. d Grayscale and e color-coded T 1 map images. f T 1 map values analysis corresponding to d. g Grayscale and h pseudocolor T 2 WI. i ΔSNR corresponding to g. j Grayscale and k color-coded T 2 map images. l T 2 map values analysis corresponding to j. Tumor regions are indicated by white dotted lines 293T, PC-12, and SH-SY5Y cells over a wide range of concentrations, exhibiting no toxicity. We also explored tissue toxicity of PDGFB-FMS using hematoxylin and eosin (H&E) analysis. After 7 days of PDGFB-FMS treatment, no obvious abnormalities in the vital organs were observed, confirming the excellent biosafety of the tissue (Fig. 7d). In addition, the biodistribution analysis revealed that the reticuloendothelial system was responsible for the majority of PDGFB-FMS accumulation in the liver (Fig. 7c). Notably, PDGFB-FMS levels in the body decreased significantly over time, indicating that it can be excreted effectively. Except for the accumulation in the liver at the early stages, PDGFB-FMS shows higher accumulation in tumors than non-targeted FMS, demonstrating excellent tumor-targeting (Fig. 7e). Further research was also conducted into the pharmacokinetics of PDGFB-FMS, FMS, and Gd-DTPA. At p.i. 10 h, the residual concentrations of FMS and Gd-DTPA were only 18.9% and 4.6%, respectively, while PDGFB-FMS was 23.8%. Moreover, PDGFB-FMS had the longest half-life in the blood (t 1/2 = 3.68 h) compared to FMS (t 1/2 = 2.88 h) and Gd-DTPA (t 1/2 = 0.97 h) (Fig. 7f-h). Notably, at p.i. 72 h, PDGFB-FMS could be completely eliminated from the body, thereby avoiding the potential risks to the body posed by long-term residual [42]. To summarize, the superior biocompatibility of PDGFB-FMS and its prompt in vivo clearance facilitate its use in tumor molecular MRI.

Conclusion
In conclusion, TME-activated "off-on" T 1 -T 2 dualmode MRI nanoswitch (PDGFB-FMS) has been successfully designed and fabricated. This novel nanoswitch structure is composed of a superparamagnetic Fe 3 O 4 core and a paramagnetic Mn-doped silica shell, in which the T 1 and T 2 signals are suppressed due to the phenomenon of distance-dependent magnetic resonance tuning. When exposed to GSH and a low pH value, the T 1 -T 2 dual-mode MRI signals are activated with significant T 1 and T 2 contrast enhancement. PDGFB-FMS has been utilized for the precise diagnosis of heterotopic and orthotopic tumors due to its increased SNR and tumor-specificity. Biosafety assessments further confirm the low toxicity of PDGFB-FMS. In addition, due to its rapid biodegradability, PDGFB-FMS can be excreted out of the body in a timely manner, posing no potential risk to the body. This study, therefore, provides a promising strategy for the development of intelligent MRI contrast agents to achieve accurate diagnosis of tumors.
Additional file 1: Figure S1. The hydrodynamic sizes of different samples. Figure S2. TG curves, XPS spectra, FT-IR spectra, and zeta potential of different samples. STEM image of FMS. Figure S3. Hydrodynamic size variation of PDGFB-FMS in different media. Figure S4. TEM images of FMS after different treatments and corresponding hydrodynamic size distribution. Figure S5. T 1 WI and T 2 WI of FMS. Figure S6. r 1 and r 2 values of FMS. Figure S7. r 1 value of free Mn 2+ ions at 3.0 T, hydrodynamic size distribution of FMS-1 to FMS-5, r 1 linear fit curves and r 2 linear fit curves of FMS-1 to FMS-5. Figure S8. N 2 adsorption-desorption isotherms and corresponding pore size distribution of FS, and FMS-1 to FMS-5. Figure  S9. T 1 -T 2 WI of FMS with different Fe 3 O 4 to Mn 2+ ratios. Figure S10. Fe content of 4T1 cells. Figure S11. CLSM observation: the internalization of PC-3 cells. Figure S12. Colocalization and particle uptake ratios of 4T1 cells. Figure S13. Representative color-coded T 1 and T 2 map images of 4T1 tumor-bearing mice and T 1 and T 2 map values analysis. Figure S14. Cytotoxicity of PDGFB-FMS against PC-3 cells. Figure S15. Blood routine examination of mice. Figure S16. The viabilities of THLE-3, 293 T, PC-12, and SH-SY5Y cells. Table S1. ∆r 1 and ∆r 2 value of FMS under different pH conditions. Table S2. ∆r 1 and ∆r 2 value of FMS under different GSH conditions. Table S3. In vitro ∆R 1 and ∆R 2 of 4T1 cells treated with PDGFB-FMS at the presence of different concentrations of GSH.

Author contributions
JQ and GZ designed this study. YZ and CZ prepared materials and carried out the corresponding characterization. WL and TS carried out the cell and the animal experiments. PW, DS, XQ, and XH also performed some animal experiments. LL and GZ are all responsible for writing the manuscript. Correspondence and requests for materials should be addressed to GT and GZ. All authors read and approved the final manuscript.

Declarations
Ethics approval and consent to participate Mice used in this study were treated in accordance with the ethics committee guidelines at the Binzhou Medical University.